KR101310750B1 - biomagnetic resonance apparatus and the measuring method of the same - Google Patents

Abstract

The present invention provides an ultra-low field nuclear magnetic resonance device and a method for measuring ultra-low field nuclear magnetic resonance. The method comprises applying a first measurement bias magnetic field corresponding to the excitation frequency of a coherent biomagnetic field occurring in association with the functional connectivity of the organs of the human body, having the same direction and different intensity as the first measurement bias magnetic field. Applying a second measurement bias magnetic field, and measuring a magnetic resonance signal generated in the human body using the magnetic field measuring means.

Description

Biomagnetic resonance apparatus and its measuring method {biomagnetic resonance apparatus and the measuring method of the same}

The present invention relates to an ultra-low field nuclear magnetic resonance method, and more particularly, to an ultra-low field nuclear magnetic resonance method in which an external excitation magnetic field is replaced with a biological vibration magnetic field.

In many heart diseases, myocardial regressive or ectopic excitation is the cause. These myocardial conduction abnormalities develop into atrial arrhythmia, tachycardia, and heart failure, which cause stroke. In addition, myocardial conduction dysfunction is a mechanism of ventricular fibrillation that causes cardiac death by cardiac arrest. In order to detect the conduction abnormality of the myocardium, the catheter electrode was inserted through the aorta or the vena cava of the thigh, and the endocardial potential was measured by changing the position. Alternatively, the measurement was performed by attaching a multi-channel electrode patch or the like to the epicardium during thoracotomy. Non-invasive methods include electrocardiograms that measure the potential by attaching multiple electrodes to the rib cage and limbs, and core diagrams that measure myocardial electrical activity using ultra-sensitive magnetic sensors such as superconducting quantum interference devices (SQUIDs) or atomic magnetometers. There is a way.

One technical problem to be solved of the present invention is to measure and image a nuclear magnetic resonance signal using a biological magnetic field.

In the ultra-low magnetic field nuclear magnetic resonance measurement method according to an embodiment of the present invention, the vibration frequency of the periodic coherent biological magnetic field generated in association with the functional connectivity of the organs of the human body is a first Larmor frequency. Applying a measurement bias magnetic field; Applying a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity; And separating the frequency of the magnetic resonance signal generated in the human body from the vibration frequency of the biological magnetic field by applying the second measurement bias magnetic field and measuring the signal by using the magnetic field measuring means.

In one embodiment of the present invention, the coherent biomagnetic field may have a component in a plane perpendicular to the first measurement bias magnetic field.

In one embodiment of the present invention, using a pre-magnetizing means for applying a pre-magnetized magnetic field to pre-magnetize the human body; And deactivating the pre-magnetized magnetic field before measurement.

In one embodiment of the present invention, the direction of the pre-magnetized magnetic field may coincide with the direction of the first measurement magnetic field.

In one embodiment of the present invention, the method further comprises scanning the intensity of the first measured bias magnetic field such that the proton magnetic resonance frequency of the first measured bias magnetic field coincides with the oscillation frequency of the coherent biomagnetic field. Can be.

In one embodiment of the present invention, the magnetic field measuring means may be a superconducting quantum interference element or an optical pumping atomic magnetometer.

In an embodiment of the present invention, the method may further include providing a gradient magnetic field to the human body.

In one embodiment of the present invention, the gradient magnetic field may include at least one of a resonance region selection gradient magnetic field, a gradient echo magnetic field, and an encoding gradient magnetic field for an image.

In one embodiment of the present invention, the resonance region selection gradient magnetic field includes at least one of a first resonance region selection gradient magnetic field, a second resonance region selection gradient magnetic field, and a third resonance region selection magnetic field, the first resonance A region selection gradient magnetic field, the second resonance region selection gradient magnetic field, and the third resonance region selection magnetic field provide a gradient magnetic field in different directions, and the resonance region selection gradient magnetic field is applied before the second measurement bias magnetic field is applied. Can be applied.

In one embodiment of the present invention, the strength of the resonance region selective gradient magnetic field is scanned such that the sum of the intensity of the resonance region selective gradient magnetic field and the intensity of the first measured magnetic field correspond to the vibration frequency of the coherent biomagnetic field. Can be.

In one embodiment of the present invention, the gradient echo magnetic field is applied after the second measurement bias magnetic field is applied, and the gradient echo magnetic field is continuously formed with the first gradient echo magnetic field and the second A gradient echo magnetic field may be included, and the first gradient echo magnetic field may be opposite to the second gradient echo magnetic field.

In one embodiment of the present invention, the encoding gradient magnetic field is applied after a second measurement bias magnetic field is applied, the encoding gradient magnetic field comprises at least one of a first encoding gradient magnetic field and a second encoding gradient magnetic field, and The encoding gradient magnetic field may perform at least one of frequency encoding or phase encoding.

In the ultra-low magnetic field nuclear magnetic resonance measurement method according to an embodiment of the present invention, a coherent biological magnetic field having a vibration frequency generated in association with a functional connection of an organ of a human body is a proton that precesses by a first measurement bias magnetic field. Selecting a resonance region to magnetically resonate with the field; And spatially imaging a selected resonance region under a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity.

In an embodiment of the present disclosure, the selecting of the resonance region may include applying a pre-magnetic field to pre-magnetize the human body; Precessing the protons by applying a spatially uniform first measurement bias magnetic field having the same direction as that of the pre-magnetized magnetic field; Applying a resonance region selective gradient magnetic field having the same direction as the first measurement bias magnetic field and having a spatial gradient; And protons precessing by the first measurement bias magnetic field or resonance region selection gradient magnetic field are magnetically resonant with the coherent biological magnetic field and excited in a predetermined space or region.

In one embodiment of the present invention, the spatial imaging may include applying a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field to the human body; Continuously applying a first gradient echo magnetic field and a second gradient echo magnetic field to form a gradient echo; Applying an encoding gradient magnetic field while the first gradient echo magnetic field is applied; And measuring a gradient echo signal while the second gradient echo magnetic field is applied.

In one embodiment of the present invention, the encoding gradient magnetic field may include at least one of a phase encoding gradient magnetic field and a frequency encoding gradient magnetic field.

Ultra-low magnetic field nuclear magnetic resonance apparatus according to an embodiment of the present invention magnetic shielding means; Magnetic field measuring means disposed in proximity to a measurement object arranged inside the magnetic shielding means; And applying a first measurement bias magnetic field corresponding to the proton magnetic resonance frequency to coincide with the vibration frequency of the coherent biological magnetic field generated in association with the functional connectivity of the organs of the human body, and continuously increasing the intensity of the first measurement bias magnetic field. And a measurement bias magnetic field generating means for altering to apply a second measurement bias magnetic field, wherein the magnetic field measuring means measures a magnetic resonance signal generated from the measurement object.

In one embodiment of the present invention, it may further include a pre-magnetization means for pre-magnetizing the measurement object.

In one embodiment of the present invention, it may further comprise a gradient magnetic field generating means for providing a gradient magnetic field to the measurement object.

In one embodiment of the present invention, the gradient magnetic field generating means comprises a resonance region selection gradient magnetic field generating means for selecting a resonance region; Gradient echo gradient magnetic field generating means for generating a gradient echo signal; And encoding gradient magnetic field generating means for generating an encoding gradient magnetic field.

In one embodiment of the present invention, the first measurement bias magnetic field may be scanned.

The method for measuring functional connectivity of the brain according to an embodiment of the present invention uses a very low-field nuclear magnetic resonance apparatus to detect functional connectivity of the brain and is resonated by a coherent biological magnetic field of a specific frequency in the brain of the human body. Magnetic resonance signals generated from protons are measured and spatially imaged.

The method for measuring functional connectivity of the brain according to an embodiment of the present invention uses a very low-field nuclear magnetic resonance apparatus to detect functional connectivity of the brain and is resonated by a coherent biological magnetic field of a specific frequency in the brain of the human body. Magnetic resonance signals generated from protons are measured and spatially imaged. The method for measuring functional connectivity of the brain may include applying a first measurement bias magnetic field having a vibration frequency of an EEG magnetic field as a Larmor frequency in association with functional connectivity of the brain; Applying a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity; And the frequency of the magnetic resonance signal generated in the brain is separated from the vibration frequency of the EEG magnetic field by the application of the second measurement bias magnetic field, and the signal is measured using magnetic field measuring means.

Nuclear magnetic resonance measurement method according to an embodiment of the present invention is non-contact, non-destructive, it can provide excellent time and space resolution. Thus, minute changes in the active current occurring inside the brain or heart can be precisely measured. Accordingly, it can be used for functional studies of brain or heart and diagnosis of functional diseases.

1 is a view illustrating a myocardial magnetic field or a biological magnetic field according to an embodiment of the present invention.
2 is a view for explaining an ultra-low magnetic field magnetic resonance device using a biological magnetic field.
3 is a view for explaining the ultra-low magnetic field nuclear magnetic resonance device of FIG.
4 is a diagram illustrating a pulse sequence of the ultra-low magnetic field nuclear magnetic resonance apparatus of FIG.
5 is a view for explaining the operation principle of the ultra-low magnetic field nuclear magnetic resonance apparatus according to an embodiment of the present invention.
6 is a diagram illustrating a situation of simulation.
Referring to FIG. 7, the 10 Hz component is a MCG signal of regressive excitation, and the 100 Hz signal is an FID signal of an HMR.
8 shows frequency selectivity of HMR.
9 is a simulation result showing the magnitude of the magnetic field in which the behavior of magnetization at each point (voxel) is measured by the squid sensor.
10A to 10D are diagrams illustrating a pulse train according to another embodiment of the present invention.
11 is a view for explaining a nuclear magnetic resonance measurement method according to another embodiment of the present invention.
12 and 13 are flowcharts illustrating a method for measuring nuclear magnetic resonance according to an embodiment of the present invention.
14 is a flowchart illustrating a method for measuring nuclear magnetic resonance according to another embodiment of the present invention.

The biological magnetic field is a magnetic signal generated from the human heart, brain, spinal cord, stomach, and the like. Can be measured with a high sensitivity magnetic sensor. Diagnostics using biomagnetic measurements can be contactless, nondestructive, and provide good time and spatial resolution. Thus, minute changes in the active current occurring inside the brain or heart can be precisely measured. Accordingly, it can be used for functional studies of brain or heart and diagnosis of functional diseases.

Initially, anatomical changes were observed to diagnose and study the living human brain. Later, localization and mapping of brain functions was attempted. Then came the study of anatomical connectivity. Currently, the importance of brain function connectivity research is emerging for the study of higher brain cognitive function. Noninvasive research equipment suitable for measurement have been developed to match the research interest of the brain.

Anatomical changes were observed using X-rays, computer tomography (CT), and magnetic resonance imaging (MRI). Functional localization was observed using magnetoencephalography (MEG), fMRI (functional MRI), and PET (positron emission tomography). Anatomical connectivity was observed using DTI (Diffusion Tensor Imaging). However, no device has yet been developed that can directly measure functional connectivity.

fMRI and PET measure biochemical information (indirect information) following cranial nerve activity, have low temporal resolution, and provide spatial discrepancies with actual neuronal activity.

Electroencephalogram (EEG) and MEG are indirect estimates of CNS location information by inverse problem solutions, and are only suitable for the analysis of cortical activity, and have a reduced sensitivity in deep neuronal activity.

Deep EEG and electrocorticogram (ECoG) can be punctured into the skull to insert electrode needles or invasive electrodes on the brain surface, and only partial information can be obtained.

Diffusion tensor imaging (DTI) images anatomical connections, not brain functions, so that spatial connections between brain functions are not directly known.

Ultra-low magnetic field magnetic resonance device according to an embodiment of the present invention measures the functional connectivity of the human organs Can provide. Ultra magnetic field magnetic resonance (BMR) devices can directly measure brain functional connectivity. Coherent brain waves that occur in association with the functional connectivity of nerves can form a coherent biomagnetic field. The coherent living magnetic field (coherent magnetic field) may be resonated with a proton. Accordingly, it is possible to directly measure the distribution mapping of non-invasive coherence of EEG and functional connectivity of the brain.

Regarding the functional connectivity of the brain, coherent brain waves can be measured in coherent EEG or spatial imaging of the coherent biomagnetic field by applying an external stimulus to or moving a part of the body. Based on this data, functional connectivity of the brain can be extracted. Based on this data, when the surgeon performs the surgery, the site of functional connectivity of the brain may not be removed. By imaging the functional connectivity of the brain, normal and abnormal individuals can be classified.

Lowering the externally measured magnetic field to uT (micro tesla) levels, the brain's nuclear magnetic resonance frequency can match the frequency of the coherent brainwave coherent biomagnetic field resulting from the collective excitation of cranial nerve currents. In addition, the functional connectivity distribution of the matched brain waves can be imaged using the MRI gradient technique. In addition, the measured signal can non-insulating the measurement bias magnetic field to remove components caused by the biological magnetic field.

The idea of direct measurement of cranial nerve current by MRI was presented by Bodurka in 1999, and in 2007, experiments were carried out by Parkes et al under high magnetic fields (greater than 1.5 T). However, the NMR signal (2 percent) due to the change in permeability of hemoglobin in high magnetic fields is large enough to mask the NMR dephasing effect (2 percent-0.002 percent) by neurocurrent. Therefore, it has been found that direct measurement of cranial nerve current is practically difficult.

The effect of changing permeability can be neglected with a low magnetic field (1-100 uT) NMR using an ultra-sensitive magnetic sensor such as SQUID. In addition, nuclear magnetic resonance can be used at several kHz, which is the frequency of generation of a pulse train of a real brain nerve. Therefore, it is possible to maximize the measurement of the NMR dephasing signal.

According to an embodiment of the present invention, instead of the cranial nerve current, the coherent electromagnetic field of coherent brain waves such as spontaneous brain waves, which are collective activities of the cranial nerves, may resonate with the nuclear magnetization of protons around the EEG generation part. By measuring this resonance signal, the distribution and connectivity of the EEG can be directly imaged. Thus, indirect and measurement limits of fMRI, MEG, etc. can be eliminated. The duration of coherence EEG is more than a few seconds, and the causuality of the functional connections of the brain can be studied with the phase difference of coherence EEG. Thus, by imaging the spatial distribution of coherent EEG, it is possible to directly explore the functional connectivity of the brain.

The measurement method according to an embodiment of the present invention may ignore all artifacts such as hemoglobin permeability change seen in Tesla level MRI as the external magnetic field drops to the uT level.

In the conventional NMR, the inductive coil is used to measure the signal, so when the external magnetic field is lowered and the frequency decreases, the induced signal decreases. However, in the measuring method according to the exemplary embodiment of the present invention, the SQUID sensor or the ultra-high sensitivity magnetic sensor of the atomic magnetometer system has no sensitivity deterioration even in the frequency band (10 Hz-several kHz) of the external magnetic field uT.

Since the signal measured at the cerebrograph is a signal several centimeters away from the signal source, the strength of the magnetic field is weak (100 fT-1pT), but the strength of the biological magnetic field applied to the proton near the location where the nerve excitement occurs is Hundreds of times stronger. Therefore, the measuring method according to an embodiment of the present invention can sufficiently cause nuclear magnetic resonance.

When the vibration frequency of the biological magnetic field resulting from the neural current coincides with the magnetic resonance frequency of the nuclear magnetization by the measurement bias magnetic field, it is possible to measure the activity of the brain nerve directly from the magnetic resonance signal.

When the biological magnetic field resulting from coherence EEG is sustained for a certain time, the phase change of the magnetic resonance signal is generated due to the effect of being added to or subtracted from the measurement bias magnetic field.

Since the relaxation time of the proton inside the human body is about 1 second, nuclear magnetization can be generated by using an additional tens of mT class electromagnets.

The cranial nerve biological magnetic field oscillating in a direction perpendicular to the measurement bias magnetic field direction causes a change in magnetization size, and a DC magnetic field in a direction parallel to the measurement bias magnetic field direction causes a change in the precession frequency of the magnetization. Therefore, it is possible to measure not only the spontaneous EEG response, which is a zero-mean oscillating brain wave, but also the DC level change of the EEG, which has been observed in the case of stroke and becomes a recent issue.

By applying an external gradient magnetic field, for example, it is possible to separate the spatial magnetic resonance frequency, thereby imaging the spatial distribution of brain nerve currents.

For example, a specific resonance region may be selected by applying an external gradient magnetic field, and imaging may be performed by using a frequency encoding method or a phase encoding method by applying another external gradient magnetic field.

Simultaneous measurement of brain magnetic resonance signal (BMR) of coherent brain waves and ultrasonography can be performed simultaneously using ultra-low magnetic field NMR. Therefore, it is possible to study the actual mechanism and connectivity of the brain by analyzing the correlation between two signals.

Electrophysiology (EP) tests examine myocardial electric activity using a catheter. The EP test inserts a catheter into the human body and measures the electrode by contacting the electrode with the endocardium. This method is invasive and always involves the risk of the procedure. In particular, the measurable site of this method is limited to the endocardium. In addition, in the case of the aorta and the vena cava, the electrode is not accessible to the opposite atrium or ventricle without puncturing the septum. In addition, the patient and physician are burdened with radiation during the procedure time to place the electrode in the correct position. Moreover, this method by itself does not give spatial information. Therefore, for the spatial mapping of myocardial electrical activity, a separate magnetic tracking device is required.

The epicardium electrode array has a great burden of thoracic surgery and requires high skill in electrode attachment, and thus cannot be used for follow-up diagnosis after the procedure.

Spatial mapping of myocardial electrical activity using electrocardiograms or magnetocardiograms is the estimation of current sources by ill-posed inverse problem solutions using noninvasive measurements. Therefore, the estimation error for a current source or a deep current source for which the constraint is not well defined is very large. Therefore, electrocardiogram or cardiac diagram is limited in clinical utilization.

Ultra-low magnetic field nuclear magnetic resonance measurement method according to an embodiment of the present invention non-invasive measurement and localization of myocardial activity that causes cardiac diseases, such as renal wave, ectopic excitation of the heart )do. Thus, the method of measurement can provide for the development of new medical devices to aid in the study, diagnosis, and treatment of heart disease.

Hereinafter, preferred embodiments of the present invention will be described in detail with reference to the accompanying drawings. However, the present invention is not limited to the embodiments described herein but may be embodied in other forms. Rather, the embodiments disclosed herein are being provided so that this disclosure will be thorough and complete, and will fully convey the concept of the invention to those skilled in the art. In the drawings, the components have been exaggerated for clarity. Portions denoted by like reference numerals denote like elements throughout the specification.

1 is a view illustrating a myocardial magnetic field or a biological magnetic field according to an embodiment of the present invention.

2 is a view for explaining an ultra-low magnetic field magnetic resonance device using a biological magnetic field.

3 is a view for explaining the ultra-low magnetic field nuclear magnetic resonance device of FIG.

4 is a diagram illustrating a pulse sequence of the ultra-low magnetic field nuclear magnetic resonance apparatus of FIG.

Referring to FIGS. 1 to 4, the method for measuring ultra-low magnetic field nuclear magnetic resonance (LMR) is based on a vibrational frequency of a coherent biological magnetic field generated in association with a functional connection of organs of a human body as a Larmor frequency. Applying a first measurement bias magnetic field, applying a second measurement bias magnetic field having the same direction and different intensity as the first measurement bias magnetic field, and measuring the frequency of the magnetic resonance signal generated in the human body Separating from the vibrational frequency of the biological magnetic field by the application of a bias magnetic field and measuring the signal using magnetic field measuring means.

The ultra-low magnetic field nuclear magnetic resonance measuring apparatus is a magnetic shield means 110, the magnetic field measuring means 160 disposed in close proximity to the measurement target 170 disposed inside the magnetic shield means 110, and the organ of the human body Applying a first measurement bias magnetic field Bm1 corresponding to the proton magnetic resonance frequency coinciding with the oscillation frequency f1 of the coherent biomagnetic field B1 occurring in association with functional connectivity, and subsequently the first measurement bias Measuring bias magnetic field generating means 140 for applying the second measuring bias magnetic field Bm2 by changing the intensity of the magnetic field Bm1. The magnetic field measuring means 160 measures a magnetic resonance signal generated from the measurement target 170.

Recurrent or ectopic excitation of the myocardium is periodic and local and repetitive. That is, the myocardium is excited with a specific frequency (fs) depending on the lesion and lesion. The myocardium in the depolarized area exhibits a potential difference from other parts of the repolarized area. The potential difference has a wave-front and generates myocardial current (I). The myocardial current I (t) generates a coherent biomagnetic field B1 or a myocardial magnetic field. The vibration frequency f1 of the myocardial magnetic field B1 is equal to the excitation frequency of myocardial electricity such as regressive excitation or ectopic excitation. The myocardial magnetic field (B1) has a strong effect on the protons forming the myocardium around the myocardial current (I). The effect decreases as the distance from the myocardial current source increases.

The myocardial magnetic field B1 at a specific vibration frequency f1 may be utilized like the B1-RF magnetic field in a general magnetic resonance image. In general magnetic resonance imaging, the B1-RF magnetic field has an RF frequency and is disposed perpendicular to the external static magnetic field, and the magnetization of the protons precessing along the external static magnetic field is tilted from the direction of the external static magnetic field by magnetic resonance.

By measuring the magnetic resonance phenomena spatially, it is possible to locate the reflexive or heterotopic excitation of the heart. In addition, the location and pathway of brain waves in the brain can be known.

On the other hand, the difference from the general magnetic resonance imaging apparatus, it uses the biological generation of the measurement bias magnetic field (Bm) and the biological magnetic field (B1) of the micro Tesla (uT) level.

     The resonance frequency due to the magnetogyric ratio of protons such as water around the excited myocardium is about 42 MHz / T. For example, suppose the frequency of the regressive wave in the paroxysmal atrial fibrillation to be found corresponds to 42 Hz. In this case, the magnitude (magnitude) of the measurement bias magnetic field (Bm), which may absorb the myocardial magnetic field or the biological magnetic field B1 of the frequency and cause magnetic resonance, corresponds to about 1 micro Tesla (uT). The biological magnetic field B1 may be formed in the z-axis direction.

Resonant protons around the myocardium that generate the biological magnetic field B1 at the vibration frequency f1 under the measurement bias magnetic field Bm may form resonance protons. In addition, non-resonant protons of the myocardium excited at frequencies other than f1 or myocardium far from the myocardium excited at frequency f1 may form non-resonant protons. The intensity of the measured bias magnetic field (Bm) is as small as one millionth of that of conventional magnetic resonance imaging (MRI). The magnitude of the measurement bias magnetic field (Bm) is smaller than the geomagnetic field of about 50 micro Tesla (uT). Therefore, in order to remove the earth's magnetic field, the measurement object 170 may be located inside the magnetic shielding means 110. The magnetic shielding means 110 may be a magnetically shielded room or an active magnetic shielding. The measurement bias magnetic field Bm may be formed on the y axis perpendicular to the direction (z axis) of the biological magnetic field.

Also, at weak measurement bias magnetic fields (Bm), alignment of the proton spins can be difficult. Therefore, the magnitude of the magnetic resonance signal actually measured is very small. Therefore, before the start of the measurement, the strong pre-polarization magnetic field Bp may be generated for a predetermined period T_p using the pre-magnetization means 150. The pre-magnetic field Bp may pre-magnetize the measurement target 170. The direction of the pre-magnetized magnetic field (Bp) may be preferably the same as the direction of the measurement bias magnetic field (Bm).

By the strong pre-magnetization magnetic field Bp, the protons are aligned in a pre-magnetization magnetic field Bp, and the measurement object 170 can be magnetized. The measurement target 170 may be a human body or an organ of the human body. The measurement object 170 may be a heart or a brain. On the other hand, the magnetic resonance precession frequency of the proton corresponding to the magnitude of the measurement bias magnetic field (Bm) is low. Thus, inductive measurement using a conventional coil whose signal grows in proportion to the frequency of the measurement signal may not provide a signal of sufficient magnitude. Accordingly, the high sensitivity magnetic field measuring means 160 may be a superconducting quantum interference element (SQUID) or an optical pumping atomic magnetometer whose measurement sensitivity is independent of the frequency of the signal.

The measurement bias magnetic field generating unit 140 generates the measurement bias magnetic field Bm and may be a conventional resistive coil. The measurement bias magnetic field generating unit 140 may be disposed in the magnetic shielding unit 110. The measurement bias magnetic field generating unit 140 may optionally scan the strength of the magnetic field. Therefore, the intensity of the measurement bias magnetic field (Bm) may be adjusted to correspond to the excitation frequency of the myocardial electricity to be measured. The measurement bias magnetic field Bm may be applied continuously or in the form of a pulse in the y-axis direction.

When the vibration frequency of the biological magnetic field B1 is 10 Hz, the frequency of the magnetic resonance signal by the first measurement bias magnetic field Bm1 is also 10 Hz. Therefore, there is a problem in that the signal of the biological magnetic field B1 due to the myocardial regressive wave (cardiac signal) is not distinguished from the signal generated in the magnetic resonance region by magnetic resonance. In order to solve this problem, the measurement bias magnetic field Bm is changed non-adiabatic.

The measurement bias magnetic field Bm may include a first measurement bias magnetic field Bm1 and a second measurement bias magnetic field Bm2. The first measurement bias magnetic field Bm1 is set to resonate with the biological magnetic field B1. The second measurement bias magnetic field Bm2 is used to change the frequency of the precession of the resonant magnetization. Accordingly, magnetic resonance occurs in the time interval T_a from when the pre-magnetization magnetic field Bp is increased to before the second measurement bias magnetic field Bm2 is applied. After the second measurement bias magnetic field Bm2 is applied, the magnetic resonance is removed and the precession frequency of magnetization is changed to correspond to the second measurement bias magnetic field Bm2. Accordingly, the signal measured at the time interval T_a includes a signal by the biological magnetic field B1 and a signal by the plane magnetization component Mxz perpendicular to the first measurement magnetic field by magnetic resonance.

However, the signal measured after the second measurement bias magnetic field Bm2 is applied may be separated from the signal by the biological magnetic field B1. The signal FID signal measured after the second measurement bias magnetic field Bm2 is applied has a second measurement bias magnetic field in which the magnetization tilted in a plane perpendicular to the first measurement bias magnetic field Bm1 by magnetic resonance Precession frequency (fp2) of (Bm2) may be due to precession motion. Accordingly, the signal FID signal measured after the second measurement bias magnetic field Bm2 is applied is a signal of the biological magnetic field B1 having a vibration frequency or magnetic resonance frequency and the second measurement bias magnetic field Bm2. The precession frequency fp2 can be separated into a magnetic resonance signal due to the precession.

The pre-magnetization means 150 may generate a pre-magnetization magnetic field Bp to pre-magnetize the measurement target 170. The pre-magnetization means 150 may enhance the nuclear magnetization of the measurement target 170 by using dynamic nuclear polarization (Dynamic Nuclear Polarization). The pre-magnetization means 150 may be a conventional resistive coil or superconducting coil. The pre-magnetization means 150 may be disposed inside the magnetic shielding means 110. In addition, the pre-magnetization means 150 may be disposed in the measurement bias magnetic field generating means 140 surrounding the measurement object 170. The prepolarization magnetic field Bp may be pulsed in the y-axis direction.

The gradient magnetic field generating unit 130 provides a gradient magnetic field to the measurement target 170. Accordingly, the nuclear magnetic resonance signal generated by the measurement target 170 may be localized. The gradient magnetic field generating means 130 may be a conventional resistive coil. The gradient magnetic field generating means 130 may be disposed between the measurement object 170 and the magnetic shielding means 110.

The gradient magnetic field BG may be in the y-axis direction. The gradient magnetic field generating means 130 is a first gradient magnetic field generating means for changing the intensity of the y-axis magnetic field (dBy / dy) along the y-axis, and a second intensity of the y-axis magnetic field (dBy / dx) for the x-axis Gradient magnetic field generating means, and third gradient magnetic field generating means in which the intensity (dBy / dz) of the y-axis magnetic field is changed along the z-axis.

The magnetic field measuring means 160 is disposed adjacent to the measurement object 170 and acquires a magnetic resonance signal emitted from the measurement object 170. The magnetic field measuring unit 160 may measure the magnetic flux in the z-axis direction. The output signal of the magnetic field measuring means 160 is provided to the measuring and analyzing unit 180. The magnetic field measuring unit 160 may measure both the biological magnetic field B1 and the magnetization component resonated by the biological magnetic field.

The magnetic field measuring unit 160 receives the output signal of the magnetic flux converter 161 for sensing and / or attenuates / amplifies the magnetic flux, and a SQUID 163 for detecting the magnetic field and converting the magnetic field into a voltage signal. And a dewar 165 for confining the refrigerant.

The SQUID 163 is a type of transducer that combines the Josephson effect of only the superconductor and the quantization of the magnetic flux to convert the change of the external magnetic flux into a voltage. The SQUID 163 is a magnetic sensor in which one or two Josephson junctions are coupled to one superconducting loop. RF SQUID can be joined by inserting one Josephson junction into a superconducting loop. DC SQUID can be joined by inserting two Josephson junctions into a superconducting loop. The RF SQUID operates in such a way as to output an alternating voltage in the RF frequency band and change in accordance with the applied magnetic flux, and the DC SQUID operates by generating a DC voltage as a function of the applied magnetic flux. It is given in the form of oscillation with a quantum value F0 (= 2.07 x 10 ^ (-15) Wb). The specific shape of the magnetic flux / voltage conversion function may be determined according to the specific structure of the DC SQUID.

The magnetic flux converter 161 detects a magnetic flux and converts the pick-up coil and / or an input coil that amplifies or attenuates it back to the SQUID 163 in the form of magnetic flux. It may include. The magnetic flux converter 161 may be formed of a superconductor. The pickup coil may have a large area in order to detect a large amount of magnetic flux. The input coil has an area similar to the SQUID to focus on the SQUID and can be wound several times to change its amplification or attenuation rate. The magnetic flux conversion unit 161 may include a magnetometer having a pickup coil composed of one loop, or a differential meter having one or more loop pairs in which pickup coils are wound in opposite directions.

The SQUID 163 may be connected to the FLL unit 188 through a conductive line. The magnetic flux converter 161 may measure the magnetic flux in the z-axis direction.

Protection is required for SQUID to operate stably under very large magnetic fields, such as a pre-magnetic field (Bp). Therefore, the ultra-low field-MRI system uses a superconducting shield 164 to protect the SQUID. However, if the entire SQUID sensor is superconducting, the SQUID cannot function as a magnetic field sensor. Therefore, when shielding using the superconductor, only the SQUID part and the input coil part of the magnetic flux conversion part are superconducting shields, and the magnetic flux conversion part 161 is placed outside the superconducting shielding. In this case, the SQUID itself is protected from the strong magnetic field by the superconducting shield 164, but the current induced from the magnetic flux converter 161 cannot be prevented from being applied to the SQUID. Therefore, in the ultra-low field-NMR system, a current limiting unit 162 may be disposed to prevent the overcurrent induced in the detection coil from being applied to the SQUID.

The measurement and analysis unit 180 may provide a frequency and a location of regressive excitation in paroxysmal atrial fibrillation using the magnetic resonance signal.

The measurement and analysis unit 180 linearizes the voltage signal of the SQUID 163 to provide a voltage signal proportional to the detected magnetic field and provides the linearized voltage signal of the FLL unit 188. And a sensor signal processor 186 for processing and removing and amplifying noise, and a sensor controller 187 for providing a control signal to the FLL unit 188.

The flux locked loop (FLL) unit 188 may include an input terminal for receiving the output signal of the SQUID 163, an integrator, a feedback linearization circuit, and a feedback coil. The FLL unit 188 may convert the amount of change in the magnetic flux into a voltage signal having a much wider range than the magnetic flux quantum value F0 and output the converted voltage signal.

The pulse sequence generator 122 receives a control signal from the controller 185 and provides a pulse sequence to the pre-magnetization coil driver 152, the measurement bias magnetic field power supply 142, and the gradient magnetic field power supply 132. do.

The magnetic field controller 101 may be synchronized with the measurement and analysis unit 180 to apply various magnetic fields to the measurement object 170. The magnetic field controller 101 may control the pre-magnetization generating means 150, the bias magnetic field generating means 140, and the gradient magnetic field generating means 130 in a series of orders.

The magnetic field controller 101 includes a pre-magnetization coil driver 152 for intermittently applying current to the pre-magnetization means 150 to form a pre-magnetization magnetic field Bp. The measurement bias magnetic field generating unit 140 that applies the measurement bias magnetic field Bm to the measurement object 170 is connected to the measurement bias magnetic field gate part 144. The measurement bias magnetic field gate portion 144 is connected to the measurement bias magnetic field power supply portion 142.

The gradient magnetic field generating means 130 is connected to the gradient magnetic field driver 134, and the gradient magnetic field driver 134 is connected to the gradient magnetic field power supply 132.

The measurement and analysis unit 180 may extract a frequency component corresponding to the second measurement bias magnetic field by processing a magnetic resonance signal (FID signal or gradient echo signal). The frequency component corresponding to the second measurement bias magnetic field may identify a resonance region resonating with the coherent biological magnetic field B1 of the measurement object.

The gradient magnetic field generating unit 130 may apply the gradient magnetic field BG to the measurement target 170. The gradient magnetic field driver 144 supplies a current to the gradient magnetic field generating unit 130 to apply the gradient magnetic field BG to the measurement target 170. The gradient magnetic field power supply 132 may provide power to the gradient magnetic field driver 134. The gradient magnetic field power supply 132 receives a pulse sequence from the pulse sequence generator 122 and supplies power to the gradient magnetic field driver 134. The gradient magnetic field generating unit 130 may generate gradient magnetic fields (dBy / dy, dBy / dz, dBy / dx).

The measurement bias magnetic field generating means 140 may generate a spatially uniform and low measurement bias magnetic field Bm. The measurement bias magnetic field generating unit 140 may be connected to the measurement bias magnetic field power supply 142. The measurement bias magnetic field gate part 144 may adjust the current applied to the measurement bias magnetic field generating means 140 to intermittently generate the measurement bias magnetic field Bm.

The pulse sequence generator 172 is provided to the pre-magnetization coil driver 152, the measurement bias magnetic field power supply 142, and the gradient magnetic field power supply 132 to generate a pulse sequence to obtain a FID signal or a gradient echo signal. can do.

The controller 165 may process the signal of the sensor signal processor 186 and control the pulse sequence generator 122 and the sensor controller 187.

Meanwhile, an optical solid state relay (SSR) may be used as a switch for turning on and off a pre-magnetizing magnetic field Bp, a measuring bias magnetic field Bm, and a gradient magnetic field BG. While the SSR is off, the pre-magnetization means 150, the measurement bias magnetic field generating means 140, and the gradient magnetic field generating means 130 may be completely shorted with the current source. The TTL signal for driving the SSR may be applied through optical communication. Accordingly, all electrical connections that may adversely affect the measurement and analysis unit 180 or the magnetic field measurement unit 160 may be removed.

According to a modified embodiment of the present invention, the first measurement bias magnetic field may be scanned to match the vibration frequency of the biological magnetic field and the resonance frequency of the first measurement bias magnetic field. The signal HMR signal acquired while the first measurement bias magnetic field is applied may increase in amplitude by magnetic resonance. In addition, the FID signal obtained after the second measurement bias magnetic field is applied may decrease with time by ending the magnetic resonance.

The HMR signal and the FID signal may be Fourier transformed into a frequency component due to a biological magnetic field and a frequency component due to magnetic resonance.

According to a modified embodiment of the present invention, conventional spatial imaging techniques may be applied so that the measurement object may be one-dimensional, two-dimensional, and three-dimensional imaging.

According to a modified embodiment of the present invention, the method for measuring functional connectivity of the brain is characterized by a coherent biological magnetic field of a specific frequency in the brain of the human body using an ultra-low field nuclear magnetic resonance apparatus to detect the functional connectivity of the brain. Magnetic resonance signals generated from resonance protons are measured and spatially imaged.

In the method of measuring functional connectivity of the brain according to a modified embodiment of the present invention, a resonance is performed by a coherent biological magnetic field of a specific frequency in the brain of the human body using an ultra-low field nuclear magnetic resonance apparatus to detect functional connectivity of the brain. Magnetic resonance signals generated from the protons are measured and spatially imaged. The method for measuring functional connectivity of the brain may include applying a first measurement bias magnetic field having a vibration frequency of an EEG magnetic field as a Larmor frequency in association with functional connectivity of the brain, having the same direction as the first measurement bias magnetic field. Applying a second measuring bias magnetic field having a different intensity, and separating the frequency of the magnetic resonance signal generated in the brain from the oscillation frequency of the EEG magnetic field by applying the second measuring bias magnetic field and dividing the signal by the magnetic field measuring means. Measure using

5 is a view for explaining the operation principle of the ultra-low magnetic field nuclear magnetic resonance apparatus according to an embodiment of the present invention.

Referring to FIG. 5, the highly sensitive magnetic field measuring unit 160 is disposed to be sensitive to a magnetic field parallel to the z-axis direction based on a Cartesian coordinate system. Both the pre-magnetic field Bp by the pre-magnetization means and the measurement bias magnetic field Bm by the measurement bias magnetic field generating means may be applied to be parallel to the y-axis direction. In this case, the nuclear spins of the protons to be measured are aligned in the y direction to form magnetization (M). The measurement starts as soon as the pre-magnetization magnetic field is turned off.

In this case, the generated magnetization M rotates about the y axis, which is the direction of the first measurement bias magnetic field Bm1. If there is no myocardial activity causing magnetic resonance, there is no magnetization component (Mz) in the z-axis in the beginning, so there is no change in the magnetic field in the z-axis even when rotating, and no signal is measured.

However, due to myocardial abnormality, a regressive wave or the like is periodically generated at the magnetic resonance frequency of the first measurement bias magnetic field Bm1, and the myocardial magnetic field B1 having the oscillation frequency f1 of the alternating current generated from the change of the heart current. When the direction is in the z-axis direction, the magnetization M aligned in the y-axis direction by the magnetic resonance phenomenon is inclined toward the x-axis or the z-side. This tilted magnetization M rotates about the y-axis, which is the direction of the first measurement bias magnetic field Bm1. The frequency of the precession is fp1. Thus, the z-axis component of the changing magnetization is made to produce a magnetic field in the z-axis direction. It is possible to measure the magnetic field in the z-axis direction by the high sensitivity magnetic field measuring means 160.

However, the vibration frequency f1 of the myocardial magnetic field B1 and the frequency fP1 of the magnetic resonance signal are the same. Therefore, there is a problem that the original myocardial regression wave overlaps with the myocardial magnetic field B1 signal and cannot be distinguished. To solve this problem, the measurement bias magnetic field is changed non-adiabatic.

That is, the first measurement bias magnetic field is changed into a second measurement bias magnetic field having different intensities. Accordingly, the frequency of the precession is changed to fp2. Thus, the magnetic resonance signal in the measured signal has fp2 and the biomagnetic field signal has fp1. Thus, magnetic resonance signals can be distinguished.

In addition, according to the direction and frequency of myocardial electric activity to be measured, it is possible to directly measure the myocardial abnormality by adjusting or scanning the applied measurement bias magnetic field (Bm1). In order to obtain spatial position information, magnetic resonance imaging (MRI) techniques using a generally known gradient magnetic field may be used.

Atrial fibrillation, a type of atrial arrhythmia, is caused by the occurrence of regressive waves due to aging or deformation of the atrial myocardium. In particular, when searching for the cause of the cause, a high frequency f wave (cyclic waveform) is found through the catheter electrode and treated through RF ablation or freezing. However, measuring the probes by contact with each other is not easy and takes a very long time. In addition, invasive examination becomes a burden even in the prognosis after the procedure.

If the configuration of the present invention is applied in this case, it is possible to image where myocardial radiofrequency (fm) occurs very safely and effectively.

The feasibility of measuring Heart Magnetic Resonance (HMR) was verified by simulation. The simulation began with constructing a myocardial reentry wave that formed the signal size of the atrial arrhythmia f-wave actually measured in the conventional cardiac studies.

6 is a diagram illustrating a situation of simulation.

Referring to FIG. 6, it is assumed that myocardial regressive waves return counterclockwise at a frequency of 10 Hz on the myocardial surface. Thus, myocardial regressive waves form coherent biocurrents. The biocurrent rotates counterclockwise in the xy plane. Accordingly, the biological magnetic field is formed in the z axis direction. The frequency of the biomagnetic field is f1. Coherent biomagnetic fields generated by myocardial regressive waves were calculated by the boundary element method using a heart and chest pain conductor model. Thus, it is possible to calculate, respectively, the biomagnetic field generated by the myocardial regressive wave at respective voxels in the volume near the myocardial regressive wave. The calculated bio magnetic field value is about 100 pT in the myocardium 1 mm away from the myocardial regressive wave. Therefore, the intensity of the RF magnetic field was sufficient to tilt the magnetization under various time constant conditions used in the experiment. By the Block equation, it is possible to calculate the dynamics of the magnetizations at each voxel by the generated biomagnetic field.

The sampling rate of the simulation was 1 kHz. The volume of the voxels used in the calculation is the space inside the cube with a 20 mm side around the myocardial regression wave. The magnetization of heart tissue was assumed to be about 30 percent of the water. The magnetization value for the calculation of the voxel was determined by comparing the measured NMR signal magnitude of water in the existing extremely low magnetic field. Pre-magnetized magnetic field size in the simulation experiment was assumed to be 200 mT. Once the pre-magnetic field is turned off, there are no strong magnetic field components that produce magnetization, so only the relaxation of magnetization occurs. At this time, T1 (spin-lattice relaxation time) and T2 (spin-spin relaxation time) were both set to 1 second. The magnetic resonance magnetic field for a tissue excited at 10 Hz is 235 nT. With the first measurement bias magnetic field Bm1 of 235 nT applied on the y axis in a direction parallel to the pre-magnetization magnetic field Bp, the magnetization of each voxel changes in the surrounding magnetic field (measured magnetic field and biological magnetic field). The magnetic field generated in response to is calculated at the position of the squid sensor.

However, since the frequency of the magnetic resonance signal generated in this case is also 10 Hz, there is a problem that the core that the original myocardial regression wave occurs also overlaps with the signal and cannot be distinguished. To solve this problem, the measurement bias magnetic field is changed non-adiabatic. That is, in this simulation, the first measurement bias magnetic field may be changed from the first measurement bias magnetic field Bm1 of 235 nT (10 Hz) to the second measurement bias magnetic field Bm2 of 2,350 nT (100 Hz) 10 times higher. .

Subsequently, a free induction decay singal (FID signal) can be measured. Accordingly, the HMR signal can be separated from the original core signal or the biomagnetic field signal. In addition, the HMR signal is limited to a particular measurement frequency (100 Hz). Therefore, by reducing the measurement bandwidth, the ease of signal measurement for system noise can be greatly increased.

Anticlockwise periodic regressive vibrations in the myocardium form oscillating biomagnetic fields that oscillate in the region of interest (ROI). When the Lamor frequency corresponding to the measurement bias magnetic field Bm coincides with the frequency of the biomagnetic field Bm, the magnetization formed by the pre-magnetization magnetic field in the region of interest begins to tilt by magnetic resonance.

In addition, the z component of the magnetization M is increased by magnetic resonance. In addition, the z component of the magnetization M is detected by a magnetic field measuring means such as SQUID. Signals detected by SQUID include biomagnetic field components due to myocardial regressive waves. Thus, in order to separate the bio magnetic field component from the detected signal, the measurement bias magnetic field is changed from the first measurement bias magnetic field to the second measurement bias magnetic field instantaneously. As a result, the magnetic resonance condition caused by the biological magnetic field is broken. As a result, the z component of the magnetization M is not satisfied by the magnetic resonance condition and is reduced by spin-spin interaction. In addition, when the intensity of the second measurement bias magnetic field Bm2 is 10 times greater than the intensity of the first measurement bias magnetic field Bm1, the frequency of precessing in the direction of the second measurement bias magnetic field Bm2 increases by 10 times. .

Referring to FIG. 7, the 10 Hz component is a MCG signal of regressive excitation, and the 100 Hz signal is an FID signal of an HMR.

8 shows frequency selectivity of HMR.

Referring to FIG. 8, the frequency selectivity of the HMR varies the intensity of the first measurement bias magnetic field Bm1, while the FID signal is measured under a fixed second measurement bias magnetic field (corresponding to 100 Hz). Accordingly, the amplitude of the 100 Hz component according to the intensity of the first measurement bias magnetic field Bm1 is displayed.

Regressive frequency selectivity has a 1 Hz resolution, and the corresponding magnetic field is about 20 nT. The nonuniformity of the magnetic field in the magnetic shield room is about 20 nT / 10 cm. Therefore, the actual measurement situation magnetic field nonuniformity is not expected to affect the frequency selectivity.

The HMR signal is a few fT in size, with both sub-fT / vHz in both SQUID and dew noise in the best-of-breed technology. According to the measuring method of the present invention, the measuring bandwidth can be limited to the line width of the signal. Therefore, assuming that the measurement bandwidth is about 0.1-1 Hz, the SNR of the HMR signal is expected to be in the order of tens. Thus, HMR signal measurement is possible.

9 is a simulation result showing the magnitude of the magnetic field in which the behavior of magnetization at each point (voxel) is measured by the squid sensor.

Referring to FIG. 9, in the simulation the squid sensor is located 2 cm above the chest surface and the myocardium is 4-6 cm deep from the chest surface. The first measurement bias magnetic field Bm1 is in the y-axis direction, and the biological magnetic field B1 is in the z-axis direction. Non-adiabatic measurement bias As the magnetization is tilted in the z-axis direction by the HMR effect, larger measurement signal components can be formed. The arrow point is the point at which the measurement bias magnetic field changes. At the time of the arrow, the measuring magnetic field changes from the first measuring magnetic field to the second measuring magnetic field, from which the generated magnetization generates a signal of 100 Hz, forming a free induction decay (FID).

The present invention can be used in a non-invasive way to detect the location of the cardiac regressive wave or ectopic excretion very accurately, it can be utilized for safe and convenient medical diagnosis. Long-term risky procedures and radiation exposure can be reduced for both patients and physicians. As a technology that can be used for diagnosis as well as prognostic observation after treatment, it can be used to develop new and innovative medical devices.

10A to 10D are diagrams illustrating a pulse train according to another embodiment of the present invention.

10A to 10D, the pre-magnetized magnetic field is applied in the y-axis direction for a predetermined time (T_BP). The pre-magnetization magnetic field may pre-magnetize the measurement object. In addition, the first measurement bias magnetic field is applied for a predetermined time. After the pre-magnetized magnetic field is removed, the biomagnetic field may provide magnetic resonance by the first measurement bias magnetic field.

In order to form a gradient echo signal, the first gradient echo magnetic field Ge1 may be applied, and the second gradient echo magnetic field Ge2 may be continuously applied. The first gradient echo magnetic field Ge1 and the second gradient echo magnetic field Ge2 may form gradient magnetic fields in opposite directions. In detail, the first gradient echo magnetic field Ge1 may form a positive dBy / dy component, and the second gradient echo magnetic field Ge2 may form a negative dBy / dy. When the second gradient echo magnetic field (Ge2) having the opposite polarity is applied at the end of the first gradient echo magnetic field (Ge1), the spindle subjected to dephasing is refocused by the gradient magnetic field having the opposite polarity. . Accordingly, a gradient echo signal may be formed.

Application of the gradient magnetic field for image acquisition in the nuclear magnetic resonance measurement method can be divided into two steps.

The ultra-low magnetic field nuclear magnetic resonance measuring method according to an embodiment of the present invention may include setting a resonance region using a gradient magnetic field, and performing spatial imaging using the gradient magnetic field. The gradient magnetic field may include a resonance region selection gradient magnetic field and an encoding gradient magnetic field.

[1] resonance area setting section (T_a)

The resonance region selection gradient magnetic fields Gr1, Gr2, and Gr3 include at least one of a first resonance region selection gradient magnetic field Gr1, a second resonance region selection gradient magnetic field Gr2, and a third resonance region selection magnetic field Gr3. can do. The first resonance region selection gradient magnetic field Gr1, the second resonance region selection gradient magnetic field Gr2, and the third resonance region selection magnetic field Gr3 may provide gradient magnetic fields in different directions. The resonance region selection gradient magnetic fields Gr1, Gr2, and Gr3 may be applied before the second measurement bias magnetic field Bm2 is applied.

The sum of the intensities of the resonance region selection gradient magnetic fields Gr1, Gr2, and Gr3 and the intensity Bm1 of the first measurement magnetic field corresponds to the vibration frequency of the coherent biological magnetic field B1, so that the resonance region selection gradient magnetic fields Gr1, The intensity of Gr2, Gr3) can be scanned.

In particular, when imaging a two-wave alpha wave (a-wave), it is necessary to know the frequency (fa) of the alpha wave and / or the location of the alpha wave generation. In order to know the frequency fa of the alpha wave, it is necessary to scan the spatially uniform first measurement bias magnetic field Bm1.

Alternatively, the resonance region selective gradient magnetic fields Gr1, Gr2, and Gr3 and the first measurement bias magnetic field Bm1 may be simultaneously applied without scanning the spatially uniform first measurement bias magnetic field Bm1. Accordingly, the magnetic field may have different intensities spatially. Accordingly, magnetic resonance may occur in a magnetic field corresponding to the frequency fa of the alpha wave in a predetermined region. The first resonance zero gradient magnetic field Gr1 may have a dBy / dy component. The second resonance-selective gradient magnetic field Gr2 may have a dBy / dx component. The third resonance zero gradient magnetic field Gr3 may have a dBy / dz component.

For example, if you image 10 Hz alpha waves of two governments, 235 nT corresponding to 10 Hz is applied to the center of the two governments, and the peripheral part is spatially inclined to be a different magnetic field so that only 10 Hz brain activity is applied to the two government centers. Let magnetic resonance occur.

Specifically, the 10 Hz alpha wave may generate a coherent biological magnetic field in the y-axis direction at the head. The first measurement bias magnetic field Bm1 and the gradient magnetic fields Gr1, Gr2, and Gr3 are applied in the y-axis direction. The magnetic field is applied in the y-axis direction so that the sum of the first measurement bias magnetic field and the gradient magnetic field is 235 nT corresponding to 10 Hz at the center of the head. As a result, magnetic resonance occurs in the two governments, and only a component where the resonance occurs lies on the x-z plane.

[2] signal spatial imaging interval (T_c)

Spatially image the distribution of resonance nuclei in which x-z components are generated.

The encoding gradient magnetic field Gen may be applied after the second measurement bias magnetic field Bm2 is applied. The encoding gradient magnetic field Gen may include at least one of the first encoding gradient magnetic field Gen1 and the second encoding gradient magnetic field Gen2. The encoding gradient magnetic field Gen may perform at least one of frequency encoding or phase encoding.

The first measurement bias magnetic field Bm1 may be changed to the second measurement bias magnetic field Bm2. Accordingly, the magnetic resonance condition is not satisfied in the two governments.

Simultaneously applying a second measuring bias magnetic field and a spatially different y-direction magnetic field (inclined magnetic field), the resonance nuclei generate signals of different frequencies depending on space. The first encoding gradient magnetic field Gen1 may have a magnitude of dBy / dx in the y-axis direction, and the second encoding gradient magnetic field Gen2 may have a magnitude of dBy / dz in the y-axis direction.

In the state of the second measurement bias magnetic field and the gradient magnetic field, the frequency of the signal has a frequency different from that of the resonance region setting section T_a which causes 10 Hz resonance (eg, 100 Hz). When the FFT of the obtained gradient echo signal, the distribution of the signal according to the frequency will be seen, and finally, since the frequency corresponds to another position, spatial imaging of the signal is possible.

11 is a view for explaining a nuclear magnetic resonance measurement method according to another embodiment of the present invention.

Referring to FIG. 11, according to a modified embodiment of the present invention, the resonant area setting step is distinguished from the existing imaging method. For example, when the spontaneous EEG current source passes, the bio magnetic field B1 in the opposite direction is formed above and below the current source.

In other words. The direction of nuclear magnetization lying down by resonance can be reversed above and below the current source. Therefore, when rotating around the measurement bias magnetic field Bm in the y direction, the phases of the magnetic fields generated are opposite to each other. Thus, the signals measured at the SQUID sensor positions cancel each other out. In practice, the magnetic resonance magnetization above the current source is closer to SQUID than below, so it can be measured slightly larger. Thus, most of the signals due to magnetic resonance magnetization are lost and only a few signals are measured.

However, if, for example, the resonance zone selection gradient magnetic field is used to spatially distinguish only the neutrons above the current source into the biomagnetic field (e.g. 0.235 nT above and 0.1 nT below), the lower neutrons will resonate. It doesn't work. Therefore, the x-z component of magnetization does not occur in the lower region. Therefore, the magnetization purely above the generation position can be measured without attenuation. If the location of the current source is unknown, the location discrimination can be performed by an iterative method.

12 and 13 are flowcharts illustrating a method for measuring nuclear magnetic resonance according to an embodiment of the present invention.

12, 13, and 10A, the nuclear magnetic resonance measuring method uses a lamor frequency as a vibrational frequency of a periodic coherent biological magnetic field generated in association with functional connectivity of organs of the human body. Applying a first measurement bias magnetic field (S130); Applying a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity (S140); And separating the frequency of the magnetic resonance signal generated in the human body from the vibration frequency of the biological magnetic field by applying the second measurement bias magnetic field and measuring the signal by using the magnetic field measuring means (S160).

The coherent biomagnetic field may have a component in a plane perpendicular to the first measurement bias magnetic field. Accordingly, the coherent biological magnetic field may perform the function of the B1-RF magnetic field in the conventional high magnetic field MRI.

The premagnetization means may be used to apply a premagnetization magnetic field to premagnetize the human body. The pre-magnetized magnetic field may then be deactivated. Accordingly, the human body or the measurement object may be magnetized beforehand. It may be preferable that the direction of the pre-magnetized magnetic field coincides with the direction of the first measurement bias magnetic field.

If the vibration frequency of the coherent biomagnetic field is unknown, the intensity of the first measurement bias magnetic field may be scanned such that the proton magnetic resonance frequency of the first measurement bias magnetic field coincides with the vibration frequency of the coherent biomagnetic field. In this case, the second measurement bias magnetic field may be constant and fixed. Accordingly, by monitoring the frequency component corresponding to the second measurement bias magnetic field, the strength of the first measurement bias magnetic field may be selected.

The magnetic field measuring means may be a superconducting quantum interference element or an optical pumping atomic magnetometer.

The gradient magnetic field may be provided to the human body (S150). The gradient magnetic field may include a resonance region selection gradient magnetic field S152, a gradient echo magnetic field S154, and an encoding gradient magnetic field S156.

A gradient echo magnetic field S154 may be applied to generate a gradient echo signal. Resonance Region Selection The gradient magnetic field S152 may be applied to select a specific resonance region. The encoding gradient magnetic field S156 can be used for two-dimensional imaging or one-dimensional imaging.

The resonance region selection gradient magnetic field may include a first resonance region selection gradient magnetic field, a second resonance region selection gradient magnetic field, and a third resonance region selection magnetic field. The first resonance region selection gradient magnetic field Gr1, the second resonance region selection gradient magnetic field Gr2, and the third resonance region selection magnetic field Gr3 may provide gradient magnetic fields in different directions. The resonance region selective gradient magnetic field may be applied before the second measurement bias magnetic field is applied. Accordingly, a specific region can be resonated in the coherent biological magnetic field by the resonance selective magnetic field.

The intensity of the resonance region selection gradient magnetic field may be scanned such that the sum of the strengths of the resonance region selection gradient magnetic field and the intensity of the first measurement magnetic field corresponds to the vibration frequency of the coherent biological magnetic field. Accordingly, the space selected can be changed.

The gradient echo magnetic field may be applied after the second measurement bias magnetic field Bm2 is applied. The gradient echo magnetic field may include a first gradient echo magnetic field Ge1 and a second gradient echo magnetic field Ge2 that are continuously formed. The first gradient echo magnetic field Ge1 may be opposite to the second gradient echo magnetic field Ge2.

The encoding gradient magnetic field may be applied after the second measurement bias magnetic field is applied. The encoding gradient magnetic field may include a first encoding gradient magnetic field and a second encoding gradient magnetic field. The encoding gradient magnetic field may perform at least one of frequency encoding or phase encoding.

The gradient echo signal may be measured after the first gradient echo magnetic field Ge1 is removed. The gradient echo signal may be Fourier transformed and signal processed. The gradient echo signal can be obtained while the FLL is operating.

14 is a flowchart illustrating a method for measuring nuclear magnetic resonance according to another embodiment of the present invention.

Referring to FIG. 14, the nuclear magnetic resonance measuring method is such that a coherent biological magnetic field having a vibrational frequency generated in association with a functional connection of an organ of a human body magnetically resonates with protons precessed by a first measurement bias magnetic field. Selecting a resonance region (S210), and spatial imaging (S220) for the selected resonance region under a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity.

Selecting the resonance region (S210) by applying a pre-magnetized magnetic field to pre-magnetize the human body (S212), by applying a spatially uniform first measurement bias magnetic field having the same direction as the direction of the pre-magnetized magnetic field Pre-washing the protons (S214), applying a resonance region selective gradient magnetic field having the same direction as the first measurement bias magnetic field and having a spatial gradient (S216), and tilting the first measurement bias magnetic field or resonance region selection Protons precessing by the magnetic field may be magnetically resonated with the coherent biological magnetic field and excited in a predetermined space or region (S218).

Spatial imaging step (S220) is the step of applying a second measurement bias magnetic field having the same direction and the different intensity to the human body in the same direction as the first measurement bias magnetic field (S222), the first gradient echo magnetic field to form a gradient echo and Continuously applying a second gradient echo magnetic field (S224), applying an encoding gradient magnetic field while the first gradient echo magnetic field is applied (S226), and a gradient echo signal while the second gradient echo magnetic field is applied It may include the step (S228) to measure. The encoding gradient magnetic field may include at least one of a phase encoding gradient magnetic field and a frequency encoding gradient magnetic field.

While the present invention has been particularly shown and described with reference to exemplary embodiments thereof, it is to be understood that the invention is not limited to the disclosed exemplary embodiments, And all of the various forms of embodiments that can be practiced without departing from the technical spirit.

110: magnetic shielding means
160: magnetic field measuring means
140: measurement bias magnetic field generating means
170: measurement target

Claims (23)

Applying a first measurement bias magnetic field whose frequency of vibration of a periodic coherent biological magnetic field that occurs in association with the functional connectivity of the organs of the human body as a Larmor frequency;
Applying a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity; And
And separating the frequency of the magnetic resonance signal generated in the human body from the vibration frequency of the biological magnetic field by applying the second measurement bias magnetic field and measuring the magnetic resonance signal using a magnetic field measuring means. Ultra-low magnetic field nuclear magnetic resonance measurement at Tesla level. The method according to claim 1,
And said coherent biomagnetic field has a component in a plane perpendicular to said first measurement bias magnetic field. The method according to claim 1,
Applying a pre-magnetic field to pre-magnetize the human body using pre-magnetization means; And
The method of claim 1, further comprising the step of inactivating the pre-magnetized magnetic field. The method of claim 3,
And the direction of the pre-magnetized magnetic field coincides with the direction of the first measurement bias magnetic field. The method according to claim 1,
And scanning the intensity of the first measured bias magnetic field such that the proton magnetic resonance frequency of the first measured bias magnetic field matches the oscillation frequency of the coherent biomagnetic field. Low-field nuclear magnetic resonance measurement method. The method according to claim 1,
And said magnetic field measuring means is a superconducting quantum interference element or an optical pumping atomic magnetometer. The method according to claim 1,
Ultrasonic field nuclear magnetic resonance measurement method of the micro Tesla level further comprising providing a gradient magnetic field to the human body. The method of claim 7, wherein
And the gradient magnetic field includes at least one of a resonance region selection gradient magnetic field, a gradient echo magnetic field, and an encoding gradient magnetic field for an image. The method of claim 8,
The resonance region selection gradient magnetic field includes at least one of a first resonance region selection gradient magnetic field, a second resonance region selection gradient magnetic field, and a third resonance region selection magnetic field,
The first resonance region selection gradient magnetic field, the second resonance region selection gradient magnetic field, and the third resonance region selection magnetic field provide gradient magnetic fields for different directions,
And the resonance region selection gradient magnetic field is applied before the second measurement bias magnetic field is applied. The method of claim 8,
The intensity of the resonance region selective gradient magnetic field is scanned such that the sum of the intensity of the resonance region selective gradient magnetic field and the intensity of the first measurement bias magnetic field correspond to the vibration frequency of the coherent biomagnetic field. Ultra-low field nuclear magnetic resonance measurement method. The method of claim 8,
The gradient echo magnetic field is applied after the second measurement bias magnetic field is applied,
The gradient echo magnetic field includes a continuously formed first gradient echo magnetic field and a second gradient echo magnetic field,
The method of claim 1, wherein the first gradient echo magnetic field is in a direction opposite to the second gradient echo magnetic field. The method of claim 8,
The encoding gradient magnetic field is applied after a second measurement bias magnetic field is applied,
The encoding gradient magnetic field comprises at least one of a first encoding gradient magnetic field and a second encoding gradient magnetic field,
And the encoding gradient magnetic field performs at least one of frequency encoding and phase encoding. Selecting a resonance region such that a coherent biological magnetic field having a vibrational frequency occurring in association with the functional connectivity of the organs of the human body magnetically resonates with protons precessed by the first measurement bias magnetic field; And
And spatially imaging a selected resonance region under a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity. The method of claim 13,
Selecting the resonance region is:
Applying a pre-magnetic field to pre-magnetize the human body;
Precessing the protons by applying a spatially uniform first measurement bias magnetic field having the same direction as that of the pre-magnetized magnetic field;
Applying a resonance region selective gradient magnetic field having the same direction as the first measurement bias magnetic field and having a spatial gradient; And
A micro tesla, wherein the protons precessing by the first measurement bias magnetic field or the resonance region selection gradient magnetic field magnetically resonate with the coherent biological magnetic field and are excited in a predetermined space or region. A method for measuring ultra-low magnetic field nuclear magnetic resonance. The method of claim 13,
Said spatial imaging comprises:
Applying to the human body a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having different intensities;
Continuously applying a first gradient echo magnetic field and a second gradient echo magnetic field to form a gradient echo;
Applying an encoding gradient magnetic field while the first gradient echo magnetic field is applied; And
And measuring at least one of the gradient echo signals while the second gradient echo magnetic field is applied. The method of claim 15,
The encoding gradient magnetic field comprises at least one of a phase encoding gradient magnetic field and a frequency encoding gradient magnetic field. Magnetic shielding means;
Magnetic field measuring means disposed in proximity to a measurement object arranged inside the magnetic shielding means; And
A first measurement bias magnetic field corresponding to the proton magnetic resonance frequency is applied to coincide with the vibration frequency of the coherent biological magnetic field generated in association with the functional connectivity of the organs of the human body, and the intensity of the first measurement bias magnetic field is continuously changed. Measuring bias magnetic field generating means for applying a second measuring bias magnetic field,
The magnetic field measuring means for measuring the magnetic resonance signal generated from the measurement target ultra-low magnetic field nuclear magnetic resonance device of the micro Tesla level. 18. The method of claim 17,
Ultra-low magnetic field nuclear magnetic resonance device of the micro Tesla level characterized in that it further comprises a pre-magnetization means for pre-magnetizing the measurement object. 18. The method of claim 17,
Ultra-low magnetic field nuclear magnetic resonance device of the micro Tesla level further comprises a gradient magnetic field generating means for providing a gradient magnetic field to the measurement object. 20. The method of claim 19,
The gradient magnetic field generating means is:
Resonance region selecting gradient magnetic field generating means for selecting a resonance region;
Gradient echo gradient magnetic field generating means for generating a gradient echo signal; And
An ultra-low magnetic field nuclear magnetic resonance apparatus according to claim 1, further comprising at least one of encoding gradient magnetic field generating means for generating an encoding gradient magnetic field. 18. The method of claim 17,
And a micro tesla level ultra-low magnetic field magnetic resonance device for scanning the first measurement bias magnetic field. In order to detect the functional connectivity of the brain, a micro-tesla-level ultra-low magnetic field magnetic resonance device is used to measure magnetic resonance signals generated from protons resonated by a coherent biological magnetic field of a specific frequency in the human brain. Method of measuring functional connectivity of the brain to spatially image. In order to detect the functional connectivity of the brain, a micro-tesla-level ultra-low magnetic field magnetic resonance device is used to measure magnetic resonance signals generated from protons resonated by a coherent biological magnetic field of a specific frequency in the human brain. In the method of measuring functional connectivity of the brain to spatially image,
Applying a first measurement bias magnetic field in which the oscillation frequency of the EEG magnetic field is a Larmor frequency in association with functional connectivity of the brain;
Applying a second measurement bias magnetic field having the same direction as the first measurement bias magnetic field and having a different intensity; And
Separating the frequency of the magnetic resonance signal generated in the brain from the vibration frequency of the EEG magnetic field by applying the second measurement bias magnetic field and measuring the magnetic resonance signal by using a magnetic field measuring means How to Measure Functional Connectivity

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